In the medical profession, the treatment of diabetes mellitus presents a continuing challenge. Diabetes is a condition characterized by abnormally high blood glucose levels. In human beings, blood glucose is primarily regulated by the pancreas, which secretes insulin. Insulin acts to facilitate the transport of glucose into the cells where it enters into various biochemical reactions.
In diabetes, the transport of glucose is impaired through defects in the function of the pancreas, allowing levels of glucose in the blood to remain high and cellular glucose levels to be markedly reduced.
There are two forms of diabetes. Type I, juvenile onset, or ketosis-prone diabetes is characterized by a reduction in the amount of insulin secreted by the pancreas. In type II, adult, or maturity onset, or ketosis-resistant diabetes, insulin secretion is normal or only minimally depressed, but the biochemical composition of the insulin is such that it facilitates glucose transport less effectively.
Since many of the complications of diabetes are believed to follow directly from the high blood levels of glucose (known as hyperglycemia), much of the clinical effort to treat diabetes has been directed toward improving glucose control by injections of exogeneous insulin.
Of the ten million diabetics in the United States, approximately 15% (all of type I diabetics and a small fraction of type II diabetics) require injections of insulin. Without some form of insulin therapy, type I diabetes is usually a fatal disease.
Glucose levels in the blood are usually dependent upon the amount of food an individual consumes. As the food is metabolized, a portion is converted to glucose and enters the bloodstream. Thus, the glucose level, and the corresponding need for an increased concentration of insulin in the bloodstream to facilitate glucose transport, vary as a function of time.
Conventional insulin therapy consists of one, and sometimes two, subcutaneous injections per day. Unfortunately, this treatment still does not provide adequate control of glucose levels in some individuals with less tolerance for deviations in glucose concentrations. For these individuals the injection of a large dose of insulin once a day increases insulin levels in the blood without any actual correspondence to the immediate concentration of glucose. For such patients there is a need to more carefully match insulin dosages to the actual glucose level in the bloodstream so that the glucose level remains within a narrower, more acceptable range.
This need is satisfied by injecting the same insulin dosage over a longer period of time, to match ambient glucose levels. This slow, constant administration of insulin is known as the baseline or basal dosage. Conversely, there is also a need to administer concentrated doses of insulin just before, or after, meals when glucose levels are highest. The concentrated dosage is known as the bolus dosage.
In developing reliable, continuous insulin infusion systems, two major approaches have been used, closed loop systems and open loop systems.
The closed loop systems typically attempt to mimic the function of the pancreas and deliver dosages of insulin proportionate and in response to augmented levels of glucose in the bloodstream. These devices incorporate a glucose sensor which continually monitors glucose levels in the blood and dispenses insulin to the patient in appropriate doses when the glucose concentration is elevated. "Closed loop" refers to a closed feedback loop which the glucose-sensing equipment forms with the patient and the insulin-dispensing equipment. Unfortunately, engineering problems, especially with the miniaturization of the glucose sensor, have to date precluded the development of a practical closed loop system.
As an alternative to the closed loop system, progress has been made in the development of a reliable open loop system. An open loop system administers dosages of insulin based upon the patient's clinical and dietary history, instead of in response to continuous measurements of glucose level in the blood.
With such a system, the physician usually fixes the basal dosage based upon the patient's previous insulin requirements. The patient usually has relatively free control over the bolus dosage administered at mealtimes. More or less insulin may be taken, depending upon the size of the meal anticipated.
To date the major obstacle to widespread usage of the open loop continuous infusion systems has been the relative unsophistication of portable infusion devices. The devices initially used for clinical studies were not designed for portable use. They were bulky, unattractive, and incorporated very few safety features.
Although much work has been done in the area of portable insulin infusion devices, none incorporates the novel features of the present invention. Typical disclosures of mechanical infusion devices include Kleinman, U.S. Pat. No. 3,964,139; Szabo, U.S. Pat. No. 3,886,938; and Whitney, U.S. Pat. No. 4,269,185. Other types of devices are disclosed by Franetski, U.S. Pat. No. 4,282,872; Blumle, U.S. Pat. No. 3,498,228; Haerten, U.S. Pat. No. 4,077,405; Buckles, U.S. Pat. No. 3,895,631; Scarlett, U.S. Pat. No. 4,274,407; Tucker, U.S. Pat. No. 4,193,397; and Ellinwood, U.S. Pat. No. 3,923,060.
Peristaltic pumps using rollers attached to a rotor have been known in the art for some time. The axes of rotation of the rollers are generally parallel to the axis of rotation of the rotor, and the perimeters of the rollers extend beyond the periphery of the rotor. A flexible tube is held around the rotor in contact with the rollers which extend from the rotor. As the rotor turns, two or more rollers alternately squeeze and release the tube as they roll across the tube's exterior surface. The elasticity of the tube walls cause it to expand to its former shape, drawing fluid into the tube behind the roller. When the next roller compresses the tube, the fluid is forced along by the squeezing of the tube, resulting in a pumping action.
Peristaltic pumps are well suited for use in medical infusion systems, because the medication fluid is isolated from direct contact with mechanical parts. This facilitates the delivery of sterile medications to the patient. In addition, such pumps are capable of delivering measured amounts of fluid, since, at every rotation, some multiple of the effective volume of the tube between each pair of rollers is delivered. Examples of such pumps include Gilmore, U.S. Pat. No. 2,668,637; Hunt, U.S. Pat. No. 3,137,240; and Muller, U.S. Pat. No. 3,384,080.
Although the peristaltic pump provides an attractive method of delivering medication to a patent, prior art pumps have not been well adapted to portable applications. The device disclosed by Muller shows a pump which is compact, but which requires extensive disassembly to replace the reservoir and delivery tube.
In applications where a pump is used to deliver medications through subcutaneous injections, it is especially important to maintain the internal areas of the delivery system in a sterile condition. This requires the tubing to be resterilized or replaced after each use. In prior art, such as Muller, the entire unit must be disassembled to remove the reservoir and delivery tubing for sterilization or replacement. If the tubing is reused too often, it will fatigue and lose the resiliency necessary for the pumping action. Additionally, in an insulin infusion system, the insulin tends to crystallize on the internal surfaces of the delivery system, thereby necessitating frequent replacement.
Thus, there has been a need for a peristaltic pump permitting ready replacement of the reservoir and delivery system to take advantage of practical low-cost injection molding technology.
Another problem frequently encountered with prior art peristaltic pumps is the difficulty in maintaining a precisely controlled pumping rate. In medical pumps, there is a need to deliver small quantities of precisely measured, and often highly concentrated, medications, over a long period of time. The injection of concentrated medications at slow pumping rates is much preferable to the injection of less concentrated medication at a faster rate, because the reservoir can be made smaller and therefore lighter. In portable pump applications, this becomes an important concern, where weight and bulk of the unit control the design criteria.
With medical infusion devices of this type, the pumping rate must be closely regulated between certain narrowly defined limits, and the diameter of the delivery tube upon which the rotors act becomes critical. A deviation in the diameter of the delivery tube causes the effective volume in the tubing between the pump rollers to vary. This makes the volume of liquid pumped for each rotation of the rotor different for tubes of different diameters. These differences in pumping rate can cause significant deviations from the desired dosage of highly-concentrated medication. Variations in tubing diameter can occur within normal tolerances for dimensions of mass-produced medical tubing.
Another causes of nonuniform pumping rates occurs as the pump operates over a period of time. The tubing upon which the pump rotor acts slowly loses some of its natural resiliancy upon which the pump depends for its operation. As this happens, the tubing flattens out, decreasing the effective volume of the tubing between the pump rotor rollers, therefore reducing pump displacement. This causes the effective pumping rate for each rotation of the pump motor to be reduced.
Still another cause of reduced pumping rate is the crystalization of insulin or other medication within the delivery tube, which also reduces the effective volume of the tube between the rotor rollers.